Computed Tomography, including both transmission (CT) and emission (PET and SPECT), is a powerful tool for non-invasively imaging anatomical structures and biological processes in small laboratory animals. With the ever-increasing number of human disease models, particularly in smaller animals such as mice and rats, the ability of high-resolution computed tomography to contribute unique information has become apparent to many researchers. Among various imaging modalities, single photon emission computed tomography (SPECT) allows the physiological study of disease models in small animals as well as in patient care. Suitable single-photon-emitting radiotracers are available for measuring a wide range of biological parameters of importance including substrate metabolism, blood flow, hypoxia, protein synthesis, and receptor characteristics. SPECT is capable of dual-isotope imaging for correlating two biological processes within a single imaging study as well as imaging trans-gene expression in vivo. In addition, radiolabeled antibodies are available that can be used to localize and characterize tumors in small animals, and they hold promise for diagnosis and treatment of cancer in humans.
The utility of SPECT has been significantly enhanced in recent years by the development of dual-modality imaging, which combines radionuclide imaging with anatomical imaging. Dual-modality imaging is typically achieved by combining SPECT images with images from x-ray CT or magnetic resonance imaging (MRI), fusing images pixel-to-pixel and simultaneously displaying selected functional and anatomical information. However, the functional and anatomical images may be geometrically inconsistent when they are obtained at different times on different pieces of equipment and with differing imaging geometries. Sequential dual-modality imaging results in increased study time, miss-registration errors, and complicated diagnostic procedures. CCD based x-ray/γ-ray detectors are being developed and have shown significant promise in overcoming these difficulties by performing near simultaneous combined SPECT/CT imaging using a single detector.
The main barriers to using SPECT or combined SPECT/CT in studies of laboratory animals have traditionally been poor spatial resolution, low sensitivity, and high cost. While most of the currently available SPECT systems are based on scintillation crystals coupled to position-sensitive photomultiplier tubes (PSPMTs), new design approaches that make use of charge coupled devices (CCDs) and position sensitive avalanche photodiodes (PSAPDs) have been shown to be effective in substantially improving the detector sensitivity and spatial resolution. Unfortunately, the current state-of-the-art scintillator technology remains the primary performance-limiting factor. A scintillator that simultaneously provides high spatial resolution, excellent stopping efficiency, high light output, fast response, and low cost is needed for small animal SPECT or combined SPECT/CT imaging.
A wide variety of new scintillators have recently become available that have characteristics that make them useful for radiation detection. Two new cerium doped halide scintillators, lanthanum chloride and lanthanum bromide (LaCl3:Ce and LaBr3:Ce) have shown potential to fulfill the requirements of scintillators used in such radiation detection methods as small animal single positron emission computed tomography (SPECT), computed tomography (CT), and combined SPECT/CT.
Crystals of LaCl3:Ce have one of the highest conversion efficiencies among known scintillators (˜50,000 photons/MeV), rapid decay time (20 ns), and a stopping efficiency comparable to that of NaI:Tl. Moreover, LaCl3:Ce shows very good linearity in energy response and has demonstrated two times better energy resolution than NaI:Tl (<3% at approximately 662 keV). The peak emission wavelength for LaCl3:Ce is λmax˜350 nm which is well matched to the quantum efficiency of photomultiplier tubes and is acceptable for new Si photodiodes (both unity gain p-i-n and avalanche Si diodes). These detectors are currently being considered for compact γ-camera and SPECT systems. The photon peak emissions are also appropriate for back thinned, ultra violet B (UVB) charge-coupled devices (CCDs) being used in SPECT/CT applications. LaCl3:Ce is therefore expected to find extensive use in nuclear medicine and digital radiology.
Crystals of LaBr3:Ce also has characteristic that are better than those of NaI:Tl. It has one of the highest conversion efficiencies among known scintillators (>63,000 photons/MeV), rapid decay time (16 ns), and a high stopping efficiency. Moreover, LaBr3:Ce shows very good linearity in energy response and has demonstrated two times better energy resolution than NaI:Tl (<3% at about 662 keV). The peak emission wavelength for LaBr3:Ce is λmax˜380 nm which like LaCl3:Ce, is well matched to the quantum efficiency of photomultiplier tubes (e.g., MAPMTs) and is acceptable for new Si photodiodes (both unity gain p-i-n and avalanche Si diodes). Further, the photon emissions are also appropriate for back thinned, UVB CCDs being used in SPECT/CT applications. As with LaCl3:Ce, LaBr3:Ce is expected to find extensive use in nuclear medicine and digital radiology.
Despite the obvious advantages of Lanthanum halide crystals, as exemplified by LaCl3:Ce and LaBr3:Ce, to date only single crystals of each have been grown and made commercially available. (Saint-Gobain Crystals and Detectors, Ohio; BrilLanCe®350 and BrilLanCe®380). Crystalline growth of these halide materials using art standard melt methods is difficult due to the stringent growth condition requirements, resulting in high production costs and limited availability. In addition, the crystals produced are usually no larger than about 5×5 cm2. Furthermore, growth from the standard melt process under equilibrium conditions tends to reject impurities from the crystal lattice, whether wanted or not, resulting in a non-uniform distribution of dopant. This causes variation in light output within the crystal and degrades the energy resolution. Finally, for these crystalline materials to be useful in high spatial resolution applications, it is necessary to produce it in a pixilated array form. This is challenging because LaHalide3:Ce is a highly hygroscopic material, making pixilation difficult and expensive. Thus, new, practical, and cost effective technologies are needed for producing structured arrays with controlled stochiometry, as are methods for protecting them from atmospheric moisture during and after fabrication.
For a CT or digital radiography detector, important performance criteria are area coverage (arrays of at least 5×5 cm2 for small animals and 20×25 cm2 or larger for human imaging); spatial resolution (better than 70 μm); dose efficiency; and speed of operation. Additionally, the scintillator used in the detector should have a rapid decay with no significant afterglow in order to minimize reconstruction artifacts due to image blurring in CT. Higher signal-to-noise ratios and dose efficiency are critical for minimizing radiation dose, and speed of operation is necessary for enhanced throughput and to resolve time-dependent phenomena. Additionally, a wide dynamic range of at least 10 bits for small animals and 16 bits for imaging humans is needed, with excellent linearity of response to dose. Thus, with its fast decay time, enhanced emission, excellent energy resolution, and high degree of response linearity, LaHalide3:Ce, could provide an appropriate scintillator if an efficient method and structure could be obtained to meet the requirements of high-resolution medical imaging.
To achieve high spatial resolution many of the current commercial SPECT systems rely on image magnification. This translates into the requirement that the detector must have a very large imaging area. For example, the NanoSPECT™ system developed by BioScan, Inc. provides reconstruction resolution on the order of 100 μm using four NaI(Tl) detectors, each consisting of a 4.0×4.0×0.5 cm3 scintillator coupled to an array of PMTs. Crystals of LaBr3:Ce or LeBr3:Ce would likely substantially improve the performance of this instrument given their superior properties, but the cost and time of production of crystals of these lanthanide halides makes the use of the materials prohibitively expensive. Current methods would require many weeks to grow crystals of an appropriate size and they would still require that the crystals be cut, shaped and polished after growth.
Traditionally, scintillation crystals coupled to photomultiplier tubes are the most common detectors in small animal imaging systems such as SPECT, PET, and gamma cameras. With the advent of new readout technologies such as pixilated avalanche photodiodes (APDs) and position-sensitive APDs, position sensitive photomultiplier tubes, and high speed CCDs, the choice of scintillator not only depends on its emission properties, but also on the spatial resolution requirements of the application. Comparisons of common inorganic scintillators used in such applications are well known in the art and can be found, for example in (U.S. Pat. No. 7,129,494). For most SPECT and gamma camera designs, NaI:Tl is the scintillator of choice at present. NaI:Tl has good light output, moderate speed, and its energy emission is well matched to photomultiplier tubes typically in use. Faster decay time and higher stopping efficiency would be desirable to achieve higher count-rates.
CsI:Tl is another common scintillator which is being used in SPECT and gamma camera designs consisting of silicon photodiodes or Si CCDs as optical sensors. In addition to excellent scintillator properties of the CsI:Tl, the fact that it can be grown in a columnar form has made it a scintillator of choice in many high resolution imaging applications such as radiography and CT. While the light output of CsI:Tl is higher than that of NaI:Tl, its decay time is even longer. The substantial afterglow associated with CsI:Tl makes it impractical for use in CT and also limits the maximum achievable count rate in some radionuclide imaging applications. Furthermore, its wavelength of emission is not very well matched to typical photomultiplier tubes used in these systems and improvement in the energy resolution of CsI:Tl is needed for some nuclear medicine imaging applications.
Other scintillators such as lutetium oxyothosilicate (LSO) and bismuth germanate (BGO) appear to be promising for PET due to their high gamma-ray stopping efficiency. However, LSO and BGO are not being considered for SPECT and/or CT due to their relatively low light output or lack of high spatial resolution which can provide practical limitations to the use of these materials when considered for pixilation. LSO is also expensive, has a radioactive component, and is not readily available in large volumes. Yttrium aluminum perovskite (YAP) has been investigated in recent designs of SPECT systems for small animals, however YAP has low light output limits, and its signal-to-noise ratio (SNR) for low energy γ-rays (e.g., 26-35 keV from 125I) can be a limitation, especially with a silicon photodiode readout.
Thus, there is an unmet need for an imaging scintillation radiation detector, which has high spatial resolution, high light output, fast response, adequate stopping efficiency, and which can provide sufficient energy resolution and surface area for small animal imaging. In order to provide high spatial resolution, the imaging scintillation radiation detector needs to be grown in a fine crystalline needle form (a microcolumnar structure), a structure which minimizes the traditional tradeoff between spatial resolution and absorption efficiency as is disclosed in this application.
Some of the established inorganic scintillators such as NaI:Tl, and CsI:Tl, which are commonly used in gamma-ray spectroscopy applications, are bright but have moderate energy resolution (˜6-7% FWHM for 662 keV photons). It is important to note that the energy resolution of these alkali-halide scintillators (and other non-alkali-halide scintillators such as LSO) is significantly worse than that expected from counting statistics (based on their light output). This issue is illustrated in FIG. 1, which plots, for a variety of alkali-halide and non-alkali-halide scintillators, the energy resolution (for 662 keV gamma-ray excitation) as a function of the mean number of photoelectrons (observed with a photomultiplier tube). The measured energy resolution of most scintillators lies considerably above the solid curve which represents the theoretical resolution based on counting statistics; the energy resolution of most scintillators is worse than expected from counting statistics.
It should also be noted that even small crystals of alkali-halide scintillators show poor energy resolution, which indicates that the degradation in energy resolution is not completely accounted for by, for example, self-absorption of light emissions and spatial non-uniformity of the dopant. The present consensus is that the main cause for degradation in the energy resolution of common scintillators, such as, for example, CsI:Tl, NaI:Tl and LSO, is non-proportionality. The luminous efficiency (i.e., the number of scintillation photons per unit energy) of the scintillator depends on the energy of the particle that excites it. A gamma-ray begins the excitation process by creating a knock-on electron by either photoelectric absorption or Compton scatter. As this primary electron traverses the scintillator, it loses energy to the scintillator (exciting it) and also produces other relatively high-energy electrons (delta-rays), which also excite the scintillator. Thus, a number of electrons will effectively excite the scintillator, even when the primary excitation source is a single gamma-ray. If the luminous efficiency is independent of the electron energy, then the number of scintillation photons produced by two γ-rays with the same energy will be the same (within counting statistics) because the sum of the electron energies is the same (and equal to the incident gamma energy). However, if the luminous efficiency depends on electron energy, then the number of scintillation photons will not necessarily be the same, and these variations can degrade the energy resolution.
Dependence of luminous efficiency on electron energy has been measured using a Compton technique, and the results for common alkali-halide scintillators are shown in FIG. 2. Ideally, the lines should be horizontal, indicating no dependence on electron energy. None of the alkali-halides possess this ideal shape, and these materials which are significantly above the theoretical curve in FIG. 1 also possess significant non-linearity (a steep slope in FIG. 2, especially at lower electron energies). Other non-alkali halide scintillators such as LSO and BGO also show strong dependence of luminous efficiency on electron energy. On the other hand, YAlO3:Ce (or YAP) shows minimal dependence between the luminous efficiency and the electron energy, which explains the agreement between its measured energy resolution and the estimated one (based on photon statistics) as shown in FIG. 1. Unfortunately, the light output of YAP is not very high.
Thus, in order to obtain a particularly useful energy resolution with scintillation crystals, it is important to have high light output, and minimal dependence between the luminous efficiency and the electron energy. Microcolumnar films of the present invention comprising a lanthanide halide, e.g., lanthanum halide (LaHalide), doped scintillator composition, including microcolumnar films of LaCl3:Ce, LaBr3:Ce, and the like; provide such materials that are capable of providing useful energy resolution. Methods for the production of these microcolumnar films of lanthanide halide scintillators are also provided.